With the aid of modern imaging methods, two- or three-dimensional image data, which can be used for visualizing an examination object that has been imaged and in addition to this can be used for further applications, are frequently generated.
The imaging methods are frequently based on the detection of X-ray radiation, with “measured projection data” being generated. For example, measured projection data can be acquired with the aid of a computer tomography system (CT system). In CT systems, a combination of X-ray source and X-ray detector that is disposed opposite the source and arranged on a gantry generally rotates round a measurement space, in which the examination object (mostly referred to hereinafter without restricting the generality as the patient). The center of rotation (also known as the “isocenter”) coincides with what is known as a system axis z. In one or a plurality of circuits, the patient is radiated with X-ray radiation from the X-ray source, with measured projection data or measured X-ray projection data being acquired with the aid of the facing X-ray detector.
The measured projection data that is generated is dependent in particular on the design of the X-ray detector. X-ray detectors usually comprise a plurality of detection units that are generally arranged in the form of a regular pixel array. The detection units each generate a detection signal for the X-ray radiation that impinges on the detection units, which signal is analyzed at specific points in time with respect to intensity and spectral distribution of the X-ray radiation in order to draw conclusions relating to the examination subject and measured projection data.
With the aid of CT imaging, for a long time it was “only” static anatomical structures that could be reproduced in image form. On the other hand, dynamic imaging and functional imaging using computer tomography were not possible for a long time, also among other things because of too high a level of radiation exposure for the patient. In the last few years, however, technological progress has meant that the possibilities for dynamic imaging have improved and have found their way into clinical routine.
What are known as dynamic CT imaging methods include images of measured CT projection data for the same examination region over a longer period of, for example, 5 to 50 s and the subsequent time-resolved reconstruction of CT image data at various times in the aforementioned time period. Essentially, two applications are common here: a first application relates to the representation of the influx and efflux of contrast agent into a vascular system. This type of dynamic imaging is also known as CT angiography. A special case of this is the “neuro DSA technique”, in which the image structures that correspond with bones can be removed from a contrast agent image and a native image by registering one on top of the other. A second application relates to representing the development over time of the contrast agent enrichment in an organ. This application is also known as dynamic CT perfusion imaging. It shows, for example, perfusion defects in the brain tissue following a stroke, makes circulatory disorders in the myocardium visible, or characterizes tumors, for example in the lung or in the abdomen, by their contrast agent enrichment and possibly follows up their response to treatment.
In order to improve the image contrast of the contrast agent used, iodine, for example, and to reduce the radiation dose for the patient, lower voltages in the X-ray tube are generally set, for dynamic CT scans, at between 70 kV and 100 kV, for instance. When such low tube voltages are selected, the mean energy level in the X-ray energy spectrum generated by an X-ray tube moves down to lower values. For example, at an X-ray tube voltage of 80 kV, the mean energy in the X-ray spectrum is around 55 keV. The X-ray energy spectrum shall be understood as the energy distribution of the X-ray quanta generated by an X-ray tube.
When reference is made hereinafter to energy in the X-ray spectrum or in the X-ray energy spectrum, this usually refers to the mean energy of the X-ray energy spectrum for the X-ray radiation that has been emitted. In the case of monoenergetic X-ray images, this mean energy can also be identical to a single energy, which corresponds in this specific case to the energy of the X-ray quanta from which the respective monoenergetic X-ray image can be reconstructed. At such low energies, we are closely approaching the K-edge of the contrast agent iodine, which occurs at an energy of 33 keV. As a result thereof, in the CT image iodine appears with higher CT values because it absorbs higher percentages of the X-ray radiation. In this way, a higher level of image noise can be allowed for a desired contrast-noise ratio and thus the radiation dose for the patient can be reduced.
It is precisely in dynamic CT imaging that a low radiation dose is required due to the long image acquisition time. However, in conventional CT systems, at low tube voltages, in particular at 80 kV, the available power for the X-ray radiation source is limited, such that, in particular in dynamic CT examinations in the trunk area of a patient, the tube current that is required for the desired contrast-noise ratio is only achieved in thinner patients and therefore only thinner patients can be examined using a low tube voltage. For stout patients on the other hand, the capacity of the X-ray sources is no longer adequate, such that increased image noise and other artifacts, such as streaks or an instability of CT values for example, appear in the CT images, which far outweigh the positive effect of the low tube voltages on the contrast/noise ratio.
If the power available to the X-ray source in the desired low tube voltages (80 kV) is not adequate with conventional dynamic CT imaging methods, then the examinations are carried out at higher tube voltages (100 kV), at which an increased power is available to the X-ray sources. However, at 100 kV the iodine contrast in the image is lower. In order to adhere to a specific contrast/noise ratio, the image noise therefore has to be suppressed more than at lower tube voltages and the radiation dose therefore has to be increased. For example, when using a tube voltage of 100 kV, a 30% higher radiation dose is required than at a tube voltage of 80 kV. Since specific dose limits have to be adhered to, the range of applications for the aforementioned imaging methods is reduced considerably.